Method and apparatus for ultrahigh sensitive optical microangiography

ABSTRACT

Embodiments herein provide an ultrahigh sensitive optical microangiography (OMAG) system that provides high sensitivity to slow flow information, such as that found in blood flow in capillaries, while also providing a relatively low data acquisition time. The system performs a plurality of fast scans (i.e., B-scans) on a fast scan axis, where each fast scan includes a plurality of A-scans. At the same time, the system performs a slow scan (i.e., C-scan), on a slow scan axis, where the slow scan includes the plurality of fast scans. A detector receives the spectral interference signal from the sample to produce a three dimensional (3D) data set. An imaging algorithm is then applied to the 3D data set in the slow scan axis to produce at least one image of the sample. In some embodiments, the imaging algorithm may separate flow information from structural information of the sample.

CROSS REFERENCE TO RELATED APPLICATIONS

The present application is a continuation of U.S. patent applicationSer. No. 13/577,857 filed Aug. 8, 2012 entitled “Method and Apparatusfor Ultrahigh Sensitive Optical Microangiography,” which is a U.S.National Phase Application of PCT/US2011/024069 entitled “Method andApparatus for Ultrahigh Sensitive Optical Microangiography,” whichclaims priority to U.S. Provisional Patent Application No. 61/302,409,filed Feb. 8, 2010, entitled “Method and Apparatus for UltrahighSensitive Optical Microangiography,” the entire disclosures of which arehereby incorporated by reference in their entirety.

The present application is related to U.S. Provisional PatentApplication No. 61/175,229, filed May 4, 2009, entitled “Method andApparatus for Quantitative Imaging of Blood Perfusion in Living Tissue”and Publication No. WO2008/039660, filed Sep. 18, 2007, entitled “InVivo Structural and Flow Imaging,” the entire disclosures of which arehereby incorporated by reference in their entirety.

GOVERNMENT INTERESTS

This invention was made with Government support under Grant/Contract No.R01HL093140, R01EB009682 and R01DC010201 awarded by the US NationalInstitute of Health. The Government has certain rights in the invention.

TECHNICAL FIELD

Embodiments herein relate to the field of imaging, and, morespecifically, to a method and apparatus for ultrahigh sensitive opticalmicroangiography.

BACKGROUND

The assessment of blood flow in living tissue provides importantinformation for diagnostics, treatment and/or management of pathologicalconditions. For example, the assessment of cutaneous (skin)microcirculations may provide important information for pathologicalconditions in dermatology, such as skin cancer, port wine staintreatment, diabetes, and plastic surgery. Similarly, assessment of theocular perfusion within the retina and choroid of the human eye isimportant in the diagnosis, treatment, and management of a number ofpathological conditions in ophthalmology, such as age-related maculardegeneration, diabetic retinopathy, and glaucoma. Accordingly, clinicaland technical tools that can noninvasively image three dimensional (3D)micro-blood vessel networks in vivo are in demand.

Several techniques have been developed to meet this need. However,current techniques suffer from various shortcomings which make themunsuitable for in vivo imaging in humans, such as low sensitivity toblood flow, insufficient resolution to provide useful depth information,and/or a long data acquisition time.

BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWINGS

Embodiments will be readily understood by the following detaileddescription in conjunction with the accompanying drawings. Embodimentsare illustrated by way of example and not by way of limitation in thefigures of the accompanying drawings.

FIG. 1 illustrates a functional block diagram of an imaging apparatus inaccordance with various embodiments of the present invention

FIG. 2 illustrates an example of a 3D data set, in accordance withvarious embodiments;

FIG. 3 illustrates an example of drive signals for an x-scanner andy-scanner, in accordance with various embodiments;

FIG. 4 illustrates another example of drive signals for an x-scanner andy-scanner, in accordance with various embodiments;

FIGS. 5A-C show images of capillaries within human dermis, taken with aultrahigh sensitive optical microangiography (UHS-OMAG) system accordingto various embodiments: (A) a Fourier domain optical coherencetomography (FDOCT) structural image; (B) a UHS-OMAG flow image; and (C)a phase resolved optical Doppler tomography (PRODT) cross-sectional flowimage from the UHS-OMAG image of FIG. 5B;

FIGS. 5D-F show images of capillaries within human dermis, taken with aprior optical microangiography (OMAG) system, for comparison with theimages in FIGS. 5A-C: (D) a Fourier domain optical coherence tomography(FDOCT) structural image; (E) an OMAG flow image; and (F) a phaseresolved optical Doppler tomography (PRODT) cross-sectional flow imagefrom the OMAG image of FIG. 5E;

FIG. 6A is a B-scan structural image of a flow phantom, and Hg. 6B is acorresponding UHS-OMAG flow image, in accordance with variousembodiments.

FIG. 7A is a velocity image obtained from a flow phantom assessed byUHS-OMAG, in accordance with various embodiments;

FIG. 7B is a plot of the velocity data across the capillary tube at theposition shown as the horizontal line in FIG. 7A, in accordance withvarious embodiments;

FIG. 7C and FIG. 7D are the corresponding results to FIGS. 7A and 7B,respectively, obtained by PRODT imaging of the same phantom, inaccordance with various embodiments;

FIG. 8A is a schematic diagram of the blood vessel system of the humanskin;

FIG. 8B is a photograph of a palm showing the scanning area;

FIG. 8C is a three dimensional (3D) rendered OMAG image of blood vesselstogether with the 3D structures, in accordance with various embodiments;

FIG. 8D is a cross sectional view of the imaged blood vessels, inaccordance with various embodiments;

FIGS. 9A-D show UHS-OMAG detailed projection views of microcirculationnetwork at different depths of skin obtained from: (A) 400-450 μm(closely representing papillary dermis); (B) 450-650 μm; (C) 650-780 μm(closely representing reticular dermis); and (D) 780-1100 μm (part ofhypodermis), respectively, in accordance with various embodiments;

FIG. 10 shows a schematic diagram of a UHS-OMAG system, in accordancewith various embodiments;

FIGS. 11A-B show in vivo UHS-OMAG imaging, in accordance with variousembodiments, of the posterior segment of eye near the macular regiontowards the optic nerve head: (A) an OMAG B-scan of microstructuresshowing morphological features, and (B) the corresponding OMAG bloodflow image;

FIGS. 11C-D show projection maps of blood flow distribution within: (C)retina and (D) choroid, obtained from one 3D scan of an area of ˜3×3 mm²near the macular region, in accordance with various embodiments; and

FIGS. 12A-F show depth-resolved images, in accordance with variousembodiments, of patent blood vessels within the retina (FIGS. 12A-C) andchoroid (FIGS. 12D-F) at the land-marked depths annotated in FIG. 11B:(A) R1—beyond 425 μm above RPE; (B) R2—between 300 and 425 μm above RPE;(C) R3—between 50 and 300 μm above RPE; (D) C1—between 0 to 70 μm belowRPE; (E) C2—between 70 to 200 μm below RPE; and (F) C3—beyond 200 μmbelow RPE.

DETAILED DESCRIPTION OF DISCLOSED EMBODIMENTS

In the following detailed description, reference is made to theaccompanying drawings which form a part hereof, and in which are shownby way of illustration embodiments that may be practiced. It is to beunderstood that other embodiments may be utilized and structural orlogical changes may be made without departing from the scope. Therefore,the following detailed description is not to be taken in a limitingsense, and the scope of embodiments is defined by the appended claimsand their equivalents.

Various operations may be described as multiple discrete operations inturn, in a manner that may be helpful in understanding embodiments;however, the order of description should not be construed to imply thatthese operations are order dependent.

The description may use perspective-based descriptions such as up/down,back/front, and top/bottom. Such descriptions are merely used tofacilitate the discussion and are not intended to restrict theapplication of disclosed embodiments.

The terms “coupled” and “connected,” along with their derivatives, maybe used. It should be understood that these terms are not intended assynonyms for each other. Rather, in particular embodiments, “connected”may be used to indicate that two or more elements are in direct physicalor electrical contact with each other. “Coupled” may mean that two ormore elements are in direct physical or electrical contact. However,“coupled” may also mean that two or more elements are not in directcontact with each other, but yet still cooperate or interact with eachother.

For the purposes of the description, a phrase in the form “NB” or in theform “A and/or B” means (A), (B), or (A and B). For the purposes of thedescription, a phrase in the form “at least one of A, B, and C” means(A), (B), (C), (A and B), (A and C), (B and C), or (A, B and C). For thepurposes of the description, a phrase in the form “(A)B” means (B) or(AB) that is, A is an optional element.

The description may use the terms “embodiment” or “embodiments,” whichmay each refer to one or more of the same or different embodiments.Furthermore, the terms “comprising,” “including,” “having,” and thelike, as used with respect to embodiments, are synonymous, and aregenerally intended as “open” terms (e.g., the term “including” should beinterpreted as “including but not limited to,” the term “having” shouldbe interpreted as “having at least,” the term “includes” should beinterpreted as “includes but is not limited to,” etc.).

With respect to the use of any plural and/or singular terms herein,those having skill in the art can translate from the plural to thesingular and/or from the singular to the plural as is appropriate to thecontext and/or application. The various singular/plural permutations maybe expressly set forth herein for sake of clarity.

In various embodiments, methods, apparatuses, and systems for ultrahighsensitive optical microangiography (UHS-OMAG) are provided. In exemplaryembodiments, a computing device may be endowed with one or morecomponents of the disclosed apparatuses and/or systems and may beemployed to perform one or more methods as disclosed herein.

Embodiments herein provide an UHS-OMAG system that delivers highsensitivity with a relatively low data acquisition time. OMAG is animaging modality that is a variation on optical coherence tomography(OCT). The imaging is based on the optical signals scattered by themoving particles. The light backscattered from a moving particle maycarry a beating frequency that may be used to distinguish scatteringsignals by the moving elements from those by the static elements.Accordingly, OMAG can be used to image the flow of particles, such asblood flow.

Various embodiments of the UHS-OMAG system may include a light source, asample arm, a reference arm and a detection arm. Illustrated in FIG. 1is an exemplary embodiment of an UHS-OMAG apparatus 100 suitable forultrahigh sensitive 2-D and 3-D flow imaging. The illustrated UHS-OMAGapparatus 100 may include some features known in the art, features whichmay not be explained in great length herein except where helpful in theunderstanding of embodiments of the present invention.

As illustrated, UHS-OMAG apparatus 100 may include a light source 10.Light source 10 may comprise any light source suitable for the purposeincluding, but not limited to, a broadband light source or a tunablelaser source. A suitable broadband light source 10 may include asuperluminescent diode. In one embodiment, light source 10 comprises asuperluminescent diode with a central wavelength of 1310 nanometers (nm)and a full-width-at-half-maximum bandwidth of 65 nm. In variousembodiments, light source 10 may be a light source having one or morelonger/shorter wavelengths, which may allow for deeper imaging. Invarious other embodiments, light source 10 may comprise a tunable lasersource such as, for example, a swept laser source.

UHS-OMAG apparatus 100 may include optics 11 to couple the light fromthe light source 10 into the system. The apparatus 100 may include abeam splitter 12 for splitting the light from the optics 11 into twobeams: a first beam provided to a reference arm 14 and a second beamprovided to a sample arm 16. In various embodiments, optics 11 mayinclude, but are not limited to, various lenses or fiber opticscomponents suitable for the purpose. Beam splitter 12 may comprise a 2×2single-mode fiber coupler or any fiber coupler suitable for the purpose.

Reference arm 14 may be configured to provide a reference light to adetection arm 30 (discussed more fully below), from the light providedby light source 10, for producing a spectral interferogram incombination with backscattered light from sample 18. Reference arm 14may include optics 20 and a mirror 22 for reflecting light from lightsource 10 for providing the reference light. Optics 20 may include, butare not limited to, various lenses suitable for the purpose.

Mirror 22 may be stationary or may be modulated. Modulation may beequivalent to frequency modulation of the detected signal at detectionarm 30. It has been observed that spectral interference signals(interferograms) may be modulated by a constant Doppler frequency by amodulated mirror 22 in the reference arm 14. The spectral interferencesignal may then be recovered by de-modulating the modulated signal atthe modulation frequency. De-modulation may be achieved using anysuitable method including, for example, a digital or opticalde-modulation method. Modulation and de-modulation of spectralinterference signals may advantageously improve the signal-to-noiseratio, resulting in an improved image quality for structural, flow, andangiographic imaging.

Sample arm 16 may be configured to provide light from light source 10 toa sample 18 by way of optics 24, a scanner 26, and optics 28. Optics 24may be used to couple the light from beam splitter 12 to scanner 26.Optics 24 may include various optical lens, for example an opticalcollimator. Scanner 26 may include a pair of x-y galvanometer scannersfor scanning sample 28 in an x-y direction. Optics 28 may comprise theappropriate optics for delivering the light from the scanner 26 ontosample 18. In various embodiments, scanner 26 may also receivebackscattered light from sample 18. Although the characteristics of thelight provided to sample 18 may depend on the particular application, insome embodiments, the lateral imaging resolution may be approximately 16micrometers (μm) determined by an objective lens that focuses light ontosample 18, with a light power on sample 18 being approximately 1milliwatt (mW).

The light returning from reference arm 14 and the light returning fromsample arm 16 (i.e., the spectral signal) may be recombined and coupledinto the beam splitter 12 for introduction to detection arm 30. Asillustrated, detection arm 30 comprises a spectrometer 34 including oneor more of various optics 36 including, but not limited to, one or morecollimators, one or more diffracting/transmission gratings, and one ormore lenses (not illustrated). In exemplary embodiments, optics 36 mayinclude a 30-millimeter (mm) focal length collimator, a 1200 lines/mmdiffracting grating, and an achromatic focusing lens with a 150 mm focallength. In various embodiments, spectrometer 34 may have a designedspectral resolution of, for example, 0.055 nm, resulting in an opticalrange of approximately 6.4 mm in aft, where the positive frequency spaceis 3.2 mm and the negative frequency space is 3.2 mm. Such parametersare exemplary and may be modified in a variety of ways in accordancewith the embodiments of the present invention.

In embodiments employing a broadband light source, spectrometer 34 mayinclude a detector such as a linear detector 38 configured to detect aspectral interference signal. Linear detector 38 may include one or moreof a line-scan camera and an area scan camera. An exemplary suitablelinear detector 38 may be a charge-coupled device (CCD).

In embodiments wherein light source 10 comprises a tunable laser ratherthan a broadband light source, however, UHS-OMAG apparatus 100 mayinclude a diffusion amplifier that may comprise one or more singleelement detectors rather than spectrometer 34. For example, one or moredual-balanced photo-diode detectors may be used.

In various embodiments, reference arm 14, sample arm 16, and detectionarm 30 may include polarization controllers (not illustrated).Polarization controllers may be configured to fine-tune the polarizationstates of light in UHS-OMAG apparatus 100. Although an UHS-OMAGapparatus within the scope of the present invention may include more orless polarization controllers, inclusion of polarization controllers inreference arm 14, sample arm 16, and detection arm 30, respectively, mayadvantageously maximize the spectral interference fringe contrast atlinear detector 38 (or another suitable detector).

In various embodiments, UHS-OMAG apparatus 100 may include one or moreuser interfaces 40 for one or more purposes including controlling lineardetector 38 and scanner 26, computing data using algorithms, displayingimages, input of data, and/or output of data, etc.

As noted above, UHS-OMAG apparatus 100 may be configured to build a 3-Ddata volume set by scanning sample 18 with a sample light in x, y, and λ(z) directions to obtain a 3-D spectral interferogram data set.

In various embodiments, the scanner 26 may include an x-scanner and ay-scanner. During the composite scan, the x-scanner may perform at leastone fast scan along a fast scan axis, and the y-scanner may perform atleast one slow scan along a slow scan axis. The fast scan axis may beorthogonal to the slow scan axis (i.e., the fast scan axis and slow scanaxis may define an x-y plane). The fast scan may also be referred toherein as a B-scan, and the fast scan axis may also be referred to asthe x-axis, the lateral axis, and/or the B-scan axis. Similarly, theslow scan may also be referred to herein as a C-scan, and the slow scanaxis may also be referred to as the y-axis, the elevational axis, and/orthe C-scan axis. Each fast scan may be performed over a fast scan timeinterval, and each slow scan may be performed over a slow scan timeinterval, where the slow scan time interval is at least twice as long asthe fast scan time interval. In some embodiments, the scanner mayperform the one or more fast scans contemporaneously with the one ormore slow scans. In such embodiments, a plurality of fast scans may beperformed during one slow scan.

In each B-scan (fast scan), there may be a number, N, of A-scans. AnA-scan may be performed in the z-axis, orthogonal to both the x-axis andthe y-axis. Each A-scan may include a number, K, of pixels, i.e., datapoints, that provide imaging depth information in the z-axis. Similarly,a C-scan (slow scan) may include a number, M, of B-scans. In variousembodiments, the numbers N, K, and M may be at least two. Accordingly,during a composite scan, a three-dimensional (3D) data set may beproduced. The 3D data set may be represented as the complex functionI(x_(i),y_(j),z_(k)), where i=1, 2, . . . N, j=1, 2, . . . M, and k=1,2, . . . , K. In various embodiments, the magnitude portion of the 3Ddata set may be represented by the scalar function A(x_(i),y_(j),z_(k)),where I(x_(i),y_(j),z_(k))=A(x_(i),y_(j),z_(k))exp(iφ).

FIG. 2 shows an example of a 3D data set 200 in accordance with variousembodiments. Data set 200 includes one C-scan 202. C-scan 202 includes anumber, M, of B-scans 204. Each B-scan 204 includes a number, N, ofA-scans 206. Each A-scan 206 includes a number, K, of pixels 208, i.e.,data points 208.

In various embodiments, an imaging algorithm may be applied to the 3Ddata set to produce at least one image. The imaging algorithm may beapplied on the slow scan axis (i.e., y-axis). In some embodiments, theimaging algorithm may separate a moving component from a structuralcomponent of the sample. The image may be a full range structural imageand/or a separated structural/flow image. In some embodiments, the imagemay be of blood flow, such as blood flow in a capillary and/or in aretina of an eye.

In this disclosure, reference is made to a prior OMAG method developedby the present applicants and described in Publication No.WO2008/039660, filed Sep. 18, 2007, entitled “In Vivo Structural andFlow Imaging,” the entire disclosure of which is hereby incorporated byreference. As compared to the prior OMAG method, the embodiments hereinprovide a higher sensitivity and a lower data acquisition time. In someembodiments, the number, N, of A-scans in a B-scan may be decreased,while the number, M, of B-scans in a C-scan may be increased comparedwith the prior OMAG method. Additionally, the imaging algorithm isapplied on the slow axis (i.e., C-scan axis, elevational axis), ratherthan on the fast axis (i.e., B-scan axis, lateral axis) as in the priorOMAG method. This provides a higher sensitivity to the moving componentof the image (i.e., can image slower speeds), while providing a lowerdata acquisition time. A high data acquisition time may be unsuitablefor in vivo use, since involuntary movement of the subject isunavoidable. Furthermore, additional imaging algorithms are providedherein that differ from the prior OMAG method, although imagingalgorithms used in the prior OMAG method may also be used in someembodiments.

In the prior OMAG method, high pass filtering is normally applied in theB scan frames (obtained in the fast scanning axis) to isolate theoptical scattering signals between the static and moving scatters. Thedetectable flow velocity, v, is thus determined by the time spacing,Δt_(A), between the adjacent A scans, i.e., v=λ/2nΔt_(A), where λ is thecentral wavelength of the light source, and n is the refractive index ofthe sample. If a red blood cell moves along the probe beam direction ata speed, v, of less than or equal to 200 μm/s (i.e., v≦200 μm/s), thenit would require a time spacing, Δt_(A), of at least about 1.5milliseconds (i.e., Δt_(A)≧˜1.5 milliseconds) for the system to samplethe moving blood cell (assuming λ=840 nm and n=1.35). This time spacingtranslates into a scanning speed of about 643 A-scans/sec. Accordingly,the total imaging time to acquire a 3D capillary flow image of a tissuevolume would be prohibitively long. The problem is exacerbated furtherunder conditions where the probe beam is substantially perpendicular tothe blow flow, such as when imaging the posterior segment of human eye,which makes the effective blood flow that is probed by the OMAG systemvery slow because the Doppler angle approaches 90 degrees.

In this disclosure, however, the algorithm is applied to the slow axis.Accordingly, the detectable flow signal is determined by the timespacing, Δt_(B), between the adjacent B scans. In one exemplaryembodiment, the B frame rate may be 300 Hz (see example one below), soΔt_(B) may be about 3.3 milliseconds. In this case, the detectable flowwould be about 140 μm/s, while the system scanning speed (i.e., A-scanrate) is not limited. Accordingly, the total imaging time to acquire a3D capillary flow image of a tissue volume would be dramaticallyreduced.

In various embodiments, the fastest flow that can be detected by theUHS-OMAG system may determined by the system imaging speed. In oneexemplary embodiment, the imaging speed may have an A-scan rate of47,000 per second (see example one below). Under this circumstance, themaximum detectable velocity may be about 30 mm/s.

In various embodiments, the slowest flow that can be detected by theUHS-OMAG system may be determined by the UHS-OMAG system phase noisefloor. In one exemplary embodiment, with the imaging speed of 47,000 persecond (see example one below), the system signal to noise ratio may beabout 85 dB. The measured slowest detectable flow velocity may be about4.0 μm/s.

In various embodiments, the x-scanner may be driven by a fast scansignal and the y-scanner may be driven by a slow scan signal. That is,the scanner may direct the probe beam along the fast scan axis and slowscan axis according to a property, such as a voltage, of the fast scansignal and the slow scan signal, respectively. The fast scan signal andslow scan signal may be any suitable waveforms, such as triangular,square, stepped, and/or sinusoidal.

Any suitable scanning protocol may be used to combine the fast scansignal and slow scan signal to produce a suitable 3D data set. Forexample, FIG. 3 shows an embodiment where the fast scan signal 302 andslow scan signal 304 are each a triangular waveform (i.e., continuousscan). The fast scan signal 302 has a higher frequency than the slowscan signal 304. As the slow scan is performed to obtain a C-scan, aplurality of fast scans are performed to produce a plurality of B-scans.In some embodiments, only one slow scan may be performed in thecomposite scan, as depicted in FIG. 3. In other embodiments, a pluralityof slow scans may be performed and combined into the composite scans.

FIG. 4 shows an alternative embodiment, in which fast scan signal 402 isa triangular waveform, and slow scan signal 404 is a stepped waveform.The slow scan signal may have a number, Q, of steps 406 to finish oneC-scan. At each step, a number, P, of B-scans may be acquired.Accordingly, the number, M, of B-scans in a C-scan may be P multipliedby Q (M=P×Q). In some embodiments, P may be at east two.

In various embodiments, any suitable imaging algorithm may be applied tothe 3D data set to produce the image of the sample, such as a blood flowimage. The imaging algorithm may be applied in the slow scan axis (i.e.,the y-direction). The imaging algorithm may be designed to extractinformation on a moving component of the sample. Applying the imagingalgorithm in the slow scan axis rather than the fast scan axis provideshigher sensitivity, thereby allowing imaging of relatively slowmovement, such as capillary blood flow within dermis which has typicalflow in the range of about 100 to 900 μm/s in a resting condition, andeven slower in diseased states.

In some embodiments, the imaging algorithm may include a differentiationoperation in the y-direction, followed by an absolute operation. In someembodiments, the differentiation and absolute operations may beperformed on the complex function I′(x_(i), y_(j), z_(k)) (i.e.,I′(x_(i),y_(j),z_(k))=|I(x_(i),y_(j),z_(k))−I(x_(i),y_(j-1),z_(k))| forj=1, 2, . . . , M). In other embodiments, the differentiation andabsolute operations may be performed on the magnitude portion A(x_(i),y_(j), z_(k)) of the complex function I(x_(i), y_(j), z_(k)) (i.e.,I′(x_(i),y_(j),z_(k))=|A(x_(i),y_(j),z_(k))−A(x_(i),y_(j-1),z_(k))| forj=1, 2, . . . , M).

In other embodiments, the imaging algorithm may include a high passfilter that may be applied to the 3D data set in the y-direction (slowscan axis) in order to separate the optical signals scattered by themoving particles from the optical signal scattered by themicrostructures, i.e., the static particles. The high pass filter mayuse any suitable type of high pass filtering function.

Furthermore, in some embodiments of the imaging algorithm, the priorOMAG method, as referenced above, may be applied to the 3D data set inthe y-direction (slow scan axis).

In embodiments where M is large compared to N, i.e., B-scans are denselytaken in the y-direction, averaging adjacent B-scans resulting from theimaging algorithm may improve the quality of final flow images. Thenumber of B-scans used to average can be any number larger than, orequal to, 2.

In embodiments where the y-scanner is driven by a stepped function (asdepicted in FIG. 4), the imaging algorithm may be applied independentlyat each step. Then, the image at each step (i.e.,I′(x_(i),y_(j),z_(k))), may be averaged to obtain a final B-scan imageof the blood flow at that step. In some embodiments, the averaging maybe done according to the following equation (Eq. (1)):

I(x _(i) ,y _(j) ,z _(k))=[Σ_(j=1) ^(p) I′(x _(i) ,y _(j) ,z _(k))]/P,where t=1, 2, . . . , Q  (1)

Accordingly, the final 3D image may include a number, Q, of B-scanimages in the y-direction.

In various embodiments, the algorithms described above may be applied tothe 3D data set in the spectral/frequency domain (i.e., the data are inwavelength (or wavenumber) format). Further, in various embodiments, alogarithm operation may be applied to the 3D data set.

In various other embodiments, the algorithms described above may beapplied to the 3D data set in the time/distance domain (i.e., the dataare in time (or distance) format). In frequency/spectral domain opticalcoherence tomography, it has been observed that the time/distance domainsignal and the spectral/frequency domain signal are a Fourier transformpair.

Example 1 Imaging Cutaneous Blood Flow at Capillary Level within Dermis

Below are a description and results of an experiment conducted using aUHS-OMAG system, in accordance with various embodiments, as applied toimaging capillary blood flows within dermis. Ideally, an imaging toolfor such an application must be able to resolve the capillary bloodflows within dermis, which are normally very slow (in the range of about100 to 900 μm/s at the resting condition, and even slower at diseasedstates). In addition, such tools must be able to provide depthinformation with an imaging resolution at a scale of capillary bloodvessel (about 10 μm). Furthermore, the imaging tool must have arelatively low data acquisition time to allow in vivo use, sinceinvoluntary movement of the subject is unavoidable.

The system setup for UHS-OMAG used in the experiment is similar to thatdescribed in R. K Wang, and L. An, “Doppler optical micro-angiographyfor volumetric imaging of vascular perfusion in vivo,” Opt. Express 17,8926-8940 (2009) (hereinafter “Article 1”), which is hereby incorporatedby reference in its entirety. Here, the main parameters are brieflydescribed. The system used a superluminescent diode as the light source,which has a central wavelength of 1310 nm and a bandwidth of 65 nm thatprovided an axial resolution in air of about 12 μm. In the sample arm, a50 mm focal length objective lens was used to achieve a lateralresolution of about 16 μm. The output light from the interferometer wasrouted to a home-built spectrometer, which had a designed spectralresolution of about 0.141 nm that provided a detectable depth range ofabout 3 mm on each side of the zero delay line. The line scan rate ofthe camera was 47,000 per second (47 kHz). With this imaging speed, thesignal to noise ratio was measured at about 85 dB with a light power onthe sample of about 3 mW.

The system applied a scanning protocol, similar to FIG. 3 (as discussedabove), that was designed to achieve ultrahigh sensitive imaging to theblood flow. First, for each B-scan (i.e. x-direction scan), 128 A-lineswere acquired with a spacing of about 15 μm between adjacent lines, thuscovering a length of about 2 mm on the tissue. The imaging rate was 300frames per second (fps). Note that with a 47 kHz line scan rate, thetheoretical imaging rate should be 367 fps. The reduced imaging rate at300 fps was due to data transfer limitations during the handshakebetween the camera and the computer.

Secondly, in the y-direction (i.e, C-scan direction), 1500 B-scans werecaptured over 2.0 mm on the tissue, which gave a spacing of about 1.3 μmbetween adjacent B-scans, equating to an oversampling factor of about 12in the C-scan direction. The whole 3D data set was captured within 5seconds.

The imaging algorithm was applied on the slow axis (C-scan direction)rather than the fast axis (B-scan direction), As discussed in Article 1,the interference signal of one B-scan captured by the CCD camera can beexpressed as the following equation (Eq. (2)):

I(t,k)=^(2S(k)E) ^(B) ^([∫) ^(−∞) ^(∞) ^(α(z,t)cos(2kn(t)z)dz+α(z) ¹^()cos [2kn(t)(z) ¹ ^(−vt)]])  (2)

where k is the wavenumber; t is the timing when an A-line was captured;E_(R) is the light reflected from the reference mirror; S(k) is thespectral density of the light source used; n is the refractive index oftissue; z is the depth coordinate; a(z, t) is the amplitude of the backscattered light; v is the velocity of moving blood cells in a bloodvessel, which is located at depth z₁. The self cross-correlation betweenthe light backscattered from different positions within the sample isnot considered in Eq. (2), because the light backscattered from thesample is weak compared to the light reflected from the referencemirror. Additionally, the DC signals are not considered, because they donot contribute to useful OMAG signals.

Prior OMAG systems used high pass filtering in the fast scanning axis,i.e. B-scan direction, to isolate the optical scattering signals betweenthe static and moving scatters. Thus, the detect-able flow velocity, v,is determined by the time spacing, Δt_(A), between the adjacent A-scans,i.e., v=λ/2nΔt_(A), wherein λ is the central wavelength of the lightsource, and n is the refractive index of the sample. If the flowvelocity in a capillary is less than or equal to 100 μm/s, then it wouldrequire Δt_(A) to be greater than or equal to 4.7 ms for the system tohave a chance to sample the blood cells flowing in the capillary. Thistime spacing translates into a scanning speed of about 213 A-scans persecond. Therefore, the total data acquisition time to acquire a 3Dcapillary flow image of a tissue volume would be prohibitively long, andnot ideal for in vivo imaging of capillary blood flows.

In order to image the slow blood flow within capillary vessels whilekeeping the data acquisition time low, the imaging algorithm isperformed in the C-scan direction (slow scan axis). In this case, Eq.(2) can still be used to represent the spectral interferogram signalcaptured by the system, except that the time variable, t, nowcorresponds to the B-scan numbers in one C-scan. With this modification,the requirement of oversampling in the B-scan direction, as was presentin the prior OMAG system, is relaxed, making it possible to have a muchfaster B-scan imaging rate, provided that the line scan camera in thespectrometer is limited or fixed. The detectable flow velocity isdetermined by the time spacing, Δt_(B), between adjacent B-scans, i.e.,v=λ/2nΔt_(B). In the system setup described here, the imaging rate is300 fps, so Δt_(B) is about 3.3 ms. Note that the imaging speed is at47,000 A scans per second, in contrast to 213 A scan per second requiredfor prior OMAG method.

In the data processing, the imaging algorithm first takes a differentialoperation on the captured B-scan spectral interferograrns along theC-scan direction, as in the following equation (Eq. (3)):

I _(flow)(t _(i) ,k)=I(t _(i) ,k),−I(t _(i-1) ,k), i=1, 2, 3 . . .1500  (3)

where i represents the index of the B-scans in the C-scan direction. Thedifferential operation suppresses the optical scattering signals fromthe static elements within scanned tissue volume. Alternatively, highpass filtering may be used. Then, a fast Fourier transform (FFT) isapplied upon every wavenumber k (t is now constant) of Eq. (2) to obtaina depth-resolved OMAG flow image with ultrahigh sensitivity to the flow.

The minimum detectable blood flow is determined by the system phasenoise floor, which can be expressed by the intensity signal to noiseratio, S, of the OMAG/OCT system by σ_(Δφ) ²=1/S. Thus, with the systemsignal to noise ratio at 85 dB, the minimum detectable flow velocitywould be about 4.0 μm/s. However, if a blood cell moves at 4 μm/s, thesystem described here would not provide a continuous trajectory for thisblood cell in the 3D OMAG flow image, i.e., the trajectory would be seenas a broken line.

Because the system is very sensitive to movement, the bulk motion of thesample may seriously degrade the final image result if the imagingalgorithm is directly applied. To solve this problem, the phasecompensation method described in Article 1 is applied to the rawinterference signal before applying the OMAG algorithm.

To test the performance of the UHS-OMAG system described above forimaging blood flow, the system was tested on the skin located on thebackside of a hand of a male volunteer. For comparison, the traditionalOMAG and phase resolved optical Doppler tomography (PRODT)cross-sectional flow images were also obtained. For these methods, thesystem captured 2000 A-scans over 2 mm at an imaging speed of 31,000A-scans per second in order to fulfill the oversampling requirement forthese previous methods.

The results are shown in FIGS. 5A-F. The images in the top row (FIGS.5A-C) are from the UHS-OMAG system while those in the bottom row (FIGS.5D-F) are from the conventional prior OMAG system. FIGS. 5A and 5D arethe Fourier domain optical coherence tomography (FDOCT) structuralimages obtained from the captured interferograms for the UHS-OMAG systemand prior OMAG system, respectively. Although they are similar, they arenot exactly the same due to the small subject movement when switchingthe system among the different approaches. However, it is sufficient toprovide a fair comparison of their ability to extract slow flowinformation. FIG. 5B shows the image from FIG. 5A after it is processedby the UHS-OMAG system. FIG. 5C shows the PRODT image of the UHS-OMAGsystem, which is based on the phase difference between adjacent B-scans.Similarly, FIG. 5E shows the image from FIG. 50 after it is processed bythe prior OMAG system. FIG. 5F shows the conventional PRODT image of theprior OMAG system, which is based on the phase difference betweenadjacent A-scans in one B-scan. In FIGS. 50 and 5F, the phasedifferences are calculated only when the structural signal is 15 dBabove the noise floor.

It is apparent that the UHS-OMAG approach outperforms the other methods.FIG. 5B shows blood flows within the papillary dermis (indicated bywhite arrows), where only capillary blood vessels are present, as wellas the blood flows within reticular dermis (indicated by red arrows),where both the capillary and larger blood vessels are present. Bycalculating the phase differences between adjacent B-scans, the bloodflow velocities within the capillaries are within the reach of theUHS-OMAG (e.g., see the white arrows in FIG. 50).

Because the conventional OMAG requires oversampling in the fast scanningdirection (i.e., B-scan), it is not sensitive to the slow blood flowwithin the capillaries, which are normally below 100 μm/s. Accordingly,the conventional PRODT approach totally failed in imaging any of bloodvessels, as seen in FIG. 5F. It should be noted that there is a globalnoise ‘flow’ background in the UHS-OMAG flow image, e.g., in FIG. 5B,which may be caused by some ‘non-moving’ scatters, such as globalmotion, etc. In this case, de-noising filters may be used to furtherenhance the UHS-OMAG flow imaging quality.

In order to test whether the flow sensitivity of UHS-OMAG approaches thesystem phase-noise floor (in this case about 4 μm/s as stared above), asecond set of images was obtained using a highly scattering flow phantomas the imaging target. The phantom was made of gelatin mixed with about1% milk to simulate the background optical heterogeneity of the tissue.In making this background tissue, precaution was taken so that the mixedgel was well solidified to minimize the possible Brownian motion ofparticles in the background. A capillary tube with an inner diameter ofabout 400 μm was submerged in this background tissue and about 2% TiO₂particle solution was flowing in it that was controlled by a precisionsyringe pump. Although such setup can precisely control the flowvelocity in the capillary tube, a flow speed as low as about 4 μm/s isdifficult to provide. This is especially true considering that if theflow is stopped, the Brownian motion of particles is unavoidable in thecapillary tube. With this experimental condition, the motion speed ofparticles due to Brownian motion would be randomly distributed within arange of several tens of microns per second. Due to these reasons, theexperiments tested the UHS-OMAG system's ability to measure the Brownianmotion of particles.

In the experiments, the capillary rube was made almost perpendicular tothe incident sample beam to avoid free fall of the scattering particleswithin the tube. The imaging results are shown in FIGS. 6A-B. FIG. 6A isthe OMAG/OCT microstructural image of the flow phantom, while FIG. 6B isthe corresponding UHS-OMAG flow image. From this result, it is clearthat UHS-OMAG is able to image the particle movements due to Brownianmotion with almost no signals detected in the background region.

To examine in more detail, the phase-resolved technique was applied tothe adjacent B-scans of the UHS-OMAG flow images to provide the velocityimage of the flow phantom above. The result is shown in FIG. 7A, whereit can be seen that the velocity values in the background region are lowwhile those within the capillary tube are contrasted out primarily dueto the Brownian motion of the particles. FIG. 7B shows a plot of thecalculated velocities across the center of the capillary tube at theposition marked as the horizontal line in FIG. 7A, where the dashed boxindicates the position of the capillary lumen. The velocity values ofparticle movements ranged from approximately −50 to 100 μm/s at thiscross-line position. The standard deviation of the values outside thedashed-box region was evaluated to be about 4.5 μm/s, which is close tothe theoretical value of about 4 μm/s. This experiment concluded thatthe UHS-OMAG system is sensitive to the flow as low as about 4 μm/s forthe system setup used in this study.

The conventional PRODT image was also obtained of the same phantom. Indoing so, the system imaging rate was set at 31,000 A-scans per second.Additionally, the A-line density across the B-scan of about 2.5 mm wasset at 4000, equating to a spacing between adjacent A-scans of about0.625 μm. The corresponding results are given in FIGS. 7C and 7D,respectively, showing that PRODT totally failed to image the Brownianmotion of the particles under the current experimental setup. Note thatthe standard deviation of velocity values shown in FIG. 7D was about 180μm/s, thus it is not surprising that PRODT is not able to achievesatisfactory imaging performance.

After evaluation of the UHS-OMAG system's imaging performance, asubsequent experiment shows its capability to image the capillary bloodflows within dermis in three dimensions. FIG. 8A shows a schematicdrawing of the blood vessel system of the human skin, in which aninterconnected network of vessels is characterized by regular structureson all levels. The human skin is composed of the cutis and the subcutis(hypodermis [HD]). The cutis is further divided into the epidermis (EP)and the dermis (DR). The dermis and subcutis are pervaded with a complexsystem of blood vessels, while the epidermis is free of vessels. Asuperficial network comprises the interface between papillary (PD) andreticular dermis (RD), while a lower network is located on the borderbetween dermis and subcutis. Vertical vessels connect both networks andthus make it complete. In the diagram, arteries are shown in red andveins in blue. To show whether the ultrahigh sensitive OMAG is able toimage the blood flow within the patent blood vessels as described above,3D blood flow images were acquired over the palm of a healthy volunteer,as shown in FIG. 8B, where the black box indicates the scanning area(about 2×2 mm²).

The 3D OMAG imaging result of blood vessel networks is shown in FIG. 8C,shown together with the 3D micro-structural image. FIG. 8D to shows across-sectional view, where the blood flows within blood vessel systemswithin the skin are clearly delineated. Because the UHS-OMAG sensitivityis as low as about 4 μm/s, even the dynamics of sweat glands are imaged.

The projection views at different land-mark depths are shown in FIGS.9A-D. FIG. 9A gives the projection view at the depths from 400 to 450μm, which corresponds to the papillary dermis where the capillaryvessels are dense (e.g., indicated by the arrows). FIG. 9B shows thedepths from 450 to 650 μm, where the vessels are almost vertical thatconnect vessel networks between papillary dermis and reticular dermis,seen as the bright spots in FIG. 9B. The blood vessel network inreticular dermis (650 to 780 μm) and hypodermis (780 to 1100 μm) areshown in FIGS. 9C and 9D, respectively. As shown, the vessel diameter issmaller in the reticular dermis than in the hypodermis. Theseobservations from ultrahigh sensitive OMAG are almost identical to thatdescribed in the literature, demonstrating the power of the ultrahighsensitive OMAG in the investigations of pathological conditions indermatology.

These experiments demonstrated an ultrahigh sensitive OMAG system toimage the volumetric microcirculation within the human skin. It wasachieved by applying the OMAG algorithm along the slow scan axis (i.e.,the C-scan direction), as opposed to the fast axis (i.e., the B scandirection). Comparing with the conventional OMAG flow image, theUHS-OMAG method delivers much better performance to extract slow flowinformation. Detailed 3D microvascular images obtained from the humanskin by the UHS-OMAG system are comparable to those described in thestandard textbook. Therefore, the ultrahigh sensitive OMAG may havegreat value in future clinical investigations of pathological conditionsin human skin.

Example 2 Imaging of Capillary Networks in Retina and Choroid of HumanEye

In another example, the UHS-OMAG system was used to obtaindepth-resolved images of the capillary networks in the retina andchoroid of the human eye. FIG. 10 shows the setup of the UHS-OMAG system1000 used to obtain the images. The UHS-OMAG system 1000 is similar tothat described in L. An, and R. K. Wang, Optics Express, 16, 11438-11452(2008) (hereinafter “Article 2”). The UHS-OMAG system 1000 used includesa light source 1002 to produce a probe beam with a wavelength centeredat 842 nm and a bandwidth of 46 nm that provided an axial resolution inair of about 8 μm. In a sample arm 1004, the light was delivered onto ahuman eye 1006 by a collimator 1008, an objective lens 1010, and anocular lens 1012. The sample arm further includes an x-scanner 1014 anda y-scanner 1016 for scanning the human eye 1006 along the x-axis andy-axis, respectively.

In a reference arm 1018, a 20 mm water cell 1020 was used to compensatefor the dispersion caused by the eye. A reference beam from the lightsource was reflected off a reference mirror 1022 to provide a referencelight. An interferogram between the reference light and the lightbackscattered from the sample was sent to a home-built ultrafastspectrometer 1024 via an optical circulator 1026. The spectrometerincluded a collimator 1027, a transmission grating 1028 (with 1200lines/mm), a camera lens 1030 with a focal length of 100 mm, and a 1024element line scan CMOS detector 1032 capable of a 140 kHz line scanrate. The spectral resolution of the designed spectrometer 1024 wasabout 0.055 nm, which provided an imaging depth of about 3 mm in air.The system sensitivity was about 90 dB measured with about 900 μW powerof light incident at the object and an exposure time of 6.9 μs with thetime interval, Δt_(A), between A-scans equal to about 7.4 μs. The system1000 further includes polarization controllers 1034, 1036, and 1038 tocontrol polarization of the light beam.

The UHS-OMAG system 1000 employed a scanning protocol, similar to FIG. 4(as discussed above), designed to achieve ultrahigh sensitive imaging ofthe flow. Firstly, the x-scanner 1012 was driven by a 400 Hz sawtoothwaveform, meaning that the imaging rate is 400 frames/sec (fps). Theduty cycle for acquiring each B-scan (i.e., x-direction scan) was about75%, in which we acquired 256 A-lines with a spacing of about 12 μmbetween adjacent lines that covered a size of about 3 mm on the retina.Secondly, the y-scanner was driven by a step function, and the entireC-scan includes 150 steps, with a spacing between adjacent steps ofabout 20 μm. In each step, 8 repeated B-scans were acquired.Accordingly, it required 3 seconds to acquire one 3D data set, coveringan area of about 3×3 mm on the retina.

An imaging algorithm was then applied on the 3D data set along the slowscan axis, i.e., C-scan direction. In this case, the detectable flowvelocity is determined by the time spacing, Δt_(B), between adjacentB-scans. Because the imaging speed was 400 fps, Δt_(B)=2.5 ms, whichwould be sufficient to image the slow flows in capillaries (see above).Finally, the calculated OMAG signals at each step were collapsed intoone B-scan through ensemble-averaging (e.g., Eq. (1)), resulting in 150B-scans to form the final C-scan image, i.e. 3D OMAG blood flowdistribution.

To demonstrate the performance of the US-OMAG system, the system wasused to obtain images on healthy volunteers. To reduce the eye and headmovement, the volunteer was asked to steer at a fixed position duringthe experiment. FIGS. 11A-D show the in vivo imaging results produced byone volume dataset captured at the macular region towards the opticnerve head. FIG. 11A shows a typical cross-sectional image (B-scan)within the OMAG structural volume, which is identical to theconventional OCT image where the typical morphological features withinretina and choroid are visualized. FIG. 11B gives the correspondingblood flow image obtained from the imaging algorithm, where thecapillary flows within the cross-section of retina are abundant(indicated by arrows), as well as the blood flow signals in the choroid.

Because of the depth-resolved nature, the blood flows in the retina canbe separated from those in the choroid. To do this, the segmentationalgorithm described in Article 2 was first used to identify the retinalpigment epithelium (RPE) layer. Then, the flow signals from the retinalvessels were identified as the OMAG flow signals 50 μm above the RPElayer (to exclude the signals from photoreceptor inner and outersegments), while those below the RPE layer were identified as thechoroidal vessels. The segmentation resulted in two volumetric flowimages, one for retina and another for choroid, which are annotated inFIG. 11B by the labels of R and C, respectively. Finally, the maximumamplitude projection (MAP) was performed on each segmented retinal andchoroidal volumes, giving the blood flow distribution maps shown inFIGS. 11C (retina) and 11D (choroid), respectively. In FIG. 11C, thereis seen a ring of blood vessels in the macular area around an avascularzone about 800 μm in diameter, denoting the fovea. This observation isin excellent agreement with the standard retinal pathology.

According to the literature, the retina consists of three layers ofcapillary networks: the radial peripapillary capillaries (RPCs, R1) andan inner (R2) and an outer layer of capillaries (R3). The RPCs are themost superficial layer of capillaries lying in the inner part of thenerve fiber layer. The inner capillaries lie in the ganglion cell layersunder and parallel to the RPCs. The outer capillary network runs fromthe inner plexiform to the outer plexiform through the inner nuclearlayers. However, the choroidal vessels in the macular region are notspecialized like those in the retina. The arteries pierce the scleraaround the optic nerve and fan out to form three vascular layers in thechoroid: inner capillary bed (near RPE layer, C1), medial arterioles andvenules (C2) and outer arteries and veins (C3). With these descriptionsas the reference, the OMAG blood flow signals obtained from the retinaland choroidal layers were further separated, with respect to the RPElayer. The segmentation is pictorially illustrated in FIG. 11B, with thelabels of R1 (about 425 μm above RPE), R2 (between 300 and 425 μm aboveRPE) and R3 (between 50 and 300 μm above RPE) for retina, and C1 (from 0to 70 μm below RPE), C2 (from 70 to 200 μm below RPE) and C3 (beyond 200μm below RPE) for choroid, respectively. After segmentation, the bloodflow MAPs within each land-marked depth are shown in FIGS. 12A-F (FIGS.12A-C for retina and FIGS. 12D-F for choroid). The results arecorrelated well with the descriptions found in the literature.

The OMAG system running at 400 fps requires about 3 seconds to acquireone 3D blood flow image representing about 3×3 mm² area on the retina.At this speed, the effects of subject movement on the final results areclearly noticeable (the horizontal lines in FIGS. 11A-D and 12A-F arethe motion artifacts). There may be several solutions to amend thisproblem, such as 1) the phase compensation algorithm developed inArticle 2 may be used to minimize the motion artifacts before theimaging algorithm is applied. In doing so, however, it would inevitablyincrease the computational load required to obtain the meaningful 3Dblood flow images because of the complexity of the compensationalgorithms, and/or 2) the imaging speed may be further increased tominimize the motion artifacts. The current system speed is limited bythe CMOS camera used in the spectrometer with a maximum line rate of 140kHz. However, a report has shown that a 4096 element CMOS camera iscapable of a line rate more than 240 kHz. Therefore, it would beexpected that if this CMOS camera is used in the system, the motionartifacts may be mitigated.

As shown above, the UHS-OMAG system may be capable of imaging detailedocular perfusion distributions within retina and choroid. Applying theimaging algorithm on the slow scanning axis, the system is sensitive tothe ocular capillary flows. In addition, due to the depth-resolvednature, the system may be able to provide detailed micro-circulationwithin different land-marked depths, the results of which are in anexcellent agreement with those described in the literature. Thedemonstrated superior imaging results show promise for future clinicalapplications for UHS-OMAG in ophthalmology.

Although certain embodiments have been illustrated and described herein,it will be appreciated by those of ordinary skill in the art that a widevariety of alternate and/or equivalent embodiments or implementationscalculated to achieve the same purposes may be substituted for theembodiments shown and described without departing from the scope. Thosewith skill in the art will readily appreciate that embodiments may beimplemented in a very wide variety of ways. This application is intendedto cover any adaptations or variations of the embodiments discussedherein. Therefore, it is manifestly intended that embodiments be limitedonly by the claims and the equivalents thereof.

1-20. (canceled)
 21. A method of imaging, comprising: receiving a threedimensional (3D) data set generated by one or more spectral interferencesignals obtained by scanning a sample with a probe beam from a lightsource, wherein the one or more spectral interference signals weregenerated by: performing a plurality of fast scans (B-scans) on a fastscan axis, each fast scan comprising a plurality of A-scans, andperforming one or more slow scans (C-scans), contemporaneously with thefast scans, on a slow scan axis, the slow scan axis being orthogonal tothe fast scan axis, and each slow scan comprising a plurality of fastscans; and applying an imaging algorithm to the 3D data set along theslow scan axis to produce at least one image of the sample.
 22. Themethod of claim 21, wherein the fast scans are driven by a firsttriangular waveform and the slow scans are driven by a second triangularwaveform, the first triangular waveform having a higher frequency thanthe second triangular waveform.
 23. The method of claim 21, wherein theslow scans are driven by a stepped waveform having a plurality of steps,and a plurality of fast scans are performed at each step of the steppedwaveform.
 24. The method of claim 21, wherein the imaging algorithm isconfigured to operate along the slow scan axis to separate a movingcomponent of the sample from a structural component of the sample. 25.The method of claim 24, wherein the imaging algorithm comprises highpass filtering along the slow scan axis in the 3D data set.
 26. Themethod of claim 21, wherein the image of the sample comprises an imageof a blood vessel network.
 27. The method of claim 21 further comprisesapplying a phase-resolved technique to the adjacent B-scans of the imageof the sample to provide blood flow velocity information.
 28. The methodof claim 21 wherein the method is applied in vivo.
 29. A system for invivo imaging, comprising: an optical micro-angiography (OMAG) apparatus;and one or more processors coupled to the OMAG apparatus and adapted tocause the OMAG apparatus to: scan a sample with a probe beam from alight source; perform a plurality of fast scans (B-scans) on a fast scanaxis, each fast scan comprising a plurality of A-scans; perform one ormore slow scans (C-scans), contemporaneously with the fast scans, on aslow scan axis, the slow scan axis being orthogonal to the fast scanaxis, and each slow scan comprising a plurality of fast scans; detectone or more spectral interference signals from the sample during thescanning to generate a three dimensional (3D) data set; and apply animaging algorithm to the 3D data set along the slow scan axis to producean image of the sample.
 30. The system of claim 29, wherein the OMAGapparatus includes an x-scanner to perform the plurality of fast scansand a y-scanner to perform the one or more slow scans, and the x-scanneris driven by a fast scan signal, and the y-scanner is driven by a slowscan signal.
 31. The system of claim 30, wherein the fast scan signalcomprises a first triangular waveform and slow scan signal comprises asecond triangular waveform, the first triangular waveform having ahigher frequency than the second triangular waveform.
 32. The system ofclaim 30, wherein the slow scan signal comprises a stepped waveformhaving a plurality of steps, and the fast scan signal is configured tocause the x-scanner to perform a plurality of fast scans during eachstep of the stepped waveform.
 33. The system of claim 29, wherein theimaging algorithm is configured to operate along the slow scan axis toseparate a moving component of the sample from a structural component ofthe sample.
 34. The system of claim 33, wherein the imaging algorithmcomprises high pass filtering along the slow scan axis in the 3D dataset.
 35. The system of claim 29, wherein the image of the samplecomprises an image of a blood vessel network.
 36. The system of claim29, further comprising applying the phase-resolved technique to theadjacent B-scans of the image of the sample to provide the blood flowvelocity information.
 37. The system of claim 29, wherein the OMAGapparatus is a Fourier domain optical coherence tomography (FD-OCT)apparatus.